WO2007023401A1 - High resolution medical imaging detector - Google Patents
High resolution medical imaging detector Download PDFInfo
- Publication number
- WO2007023401A1 WO2007023401A1 PCT/IB2006/052584 IB2006052584W WO2007023401A1 WO 2007023401 A1 WO2007023401 A1 WO 2007023401A1 IB 2006052584 W IB2006052584 W IB 2006052584W WO 2007023401 A1 WO2007023401 A1 WO 2007023401A1
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- WIPO (PCT)
- Prior art keywords
- detector
- scintillator
- sipms
- layer
- medical imaging
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Classifications
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- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/16—Measuring radiation intensity
- G01T1/161—Applications in the field of nuclear medicine, e.g. in vivo counting
- G01T1/164—Scintigraphy
- G01T1/1641—Static instruments for imaging the distribution of radioactivity in one or two dimensions using one or several scintillating elements; Radio-isotope cameras
- G01T1/1642—Static instruments for imaging the distribution of radioactivity in one or two dimensions using one or several scintillating elements; Radio-isotope cameras using a scintillation crystal and position sensing photodetector arrays, e.g. ANGER cameras
Definitions
- the present invention relates to detectors for medical imaging equipment, such as single photon emission computer tomography (SPECT) or Positron Emission Tomography (PET).
- SPECT single photon emission computer tomography
- PET Positron Emission Tomography
- FIG 1 A typical SPECT detector arrangement is shown in Figure 1 .
- Photons P such as gamma photons in the case of SPECT, enter the detector and strike a scintillator crystal X.
- the scintillator crystal X is a solid block of sodium iodide.
- the photo multiplier tube PMT converts the flash of light into electrons, which are subsequently processed to produce an image by the imaging system electronics.
- a collimator C can be used to limit the photons entering the detector to those coming at a specific orientation, while the light guide LG is used to spread the light emitted after absorption of a single gamma photon.
- Standard SPECT systems use Anger cameras, consisting of NaI crystal X and a square or hexagonal array of photo multiplier tubes PMT.
- An Anger camera has an active area of roughly 40X50 cm 2 .
- the scintillator plates are generally larger than the active area of the Anger camera.
- the intrinsic spatial resolution of an Anger camera is determined by Anger logic, a weighting algorithm that determines the point of interaction for a single gamma quantum as a function of the signal measured in some neighboring photo multiplier tubes PMT.
- the exact spatial resolution of the Anger camera depends on the size of the photo multiplier tubes PMT used. Typical photo multiplier tubes PMT have a diameter between 38 and 76 mm. As such, it is desirable to provide a medical imaging detector that solves one or more of these problems.
- the present invention is directed to an improved detector arrangement.
- a detector arrangement providing imaging information at the edge of the scintillator is provided.
- the detector arrangement provides complete information and improved spatial resolution.
- SiPMs are used to provide the geometrical coverage of the scintillator and improved spatial resolution.
- the spatial resolution can be under 2 mm.
- layers of SiPMs are placed on a front plane, and a back plane to increase the amount of information obtained from the scintillator.
- Some embodiments include a layer of SiPMs on the sides of the scintillator. Multiple layers of SiPMs improve the depth of interaction resolution of the detector.
- the overall thickness of the detector can be substantially reduced.
- Some embodiments have a combined thickness of the scintillator, any light guides included, and any layers of SiPMs that is less than 20 mm. Such embodiments make the detector much more sleek and lightweight.
- Figure 1 illustrates a prior art detector arrangement
- Figure 2 shows an illustrative example of an embodiment of the detector arrangement of the present invention.
- Figure 3 is a graphical illustration of detector size versus spatial resolution.
- Figure 4 shows an alternative embodiment of the detector arrangement of the present invention.
- Figure 5 illustrates a prior art detector arrangement with substantial non-imaged areas.
- Figures 6A and 6B illustrate a detector arrangement that provides negligible non-imaged areas.
- Figure 7 illustrates the detector arrangement shown in Figure 6A, as applied to a whole-body scan.
- Figure 8A is a prior art detector arrangement shown in application to a patient.
- Figure 8B is the detector arrangement shown in Figure 6B as applied to a patient.
- the medical imaging detector disclosed herein provides an improved spatial resolution and thereby provides improved image quality.
- the detector provides for a more compact arrangement thereby making efficient use of space.
- small avalanche photodiode cells operated in limited Ceiger mode such as silicon photomultipliers (SiPMs) are used in place of the photo multiplier tubes.
- SiPMs silicon photomultipliers
- FIG. 2 shows an illustrative embodiment of the present invention.
- the detector 10 includes an array of SiPMs 20 and a scintillator 30.
- the detector may also include one or more light guides 35 and/or a collimator 40.
- Gamma rays 44 enter the detector 10 through the collimator 40 and strike the scintillator crystal 30.
- a gamma ray strikes the scintillator, a burst of light is generated. The light is then detected by the SiPMs, which produce electrical signals that are produced into an image.
- the scintillator crystal 30 can be sodium iodide or any other scintillating material, such as, for example, cesium iodide, lanthanum bromide, lanthanum chloride, lutetium oxyorthosilicate, lutetium yttrium orthosilicate, lutetium pyrosilicate, bismuth germinate, gadolinium orthosilicate, lutetium gadolinium orthosilicate, or other suitable material.
- This invention should not be restricted in any way with regard to the scintillator, as any scintillator with sufficient light amplitude will suffice.
- the SiPMs 20 are positioned on the front plane 46 of the crystal 30. In some embodiments, SiPMs 20 are also positioned on either the sides 47 of the crystal 30, the back plane 48 of the crystal 30, or both.
- the relative small nature of the SiPMs remove the missing data caused by the gaps in the PMT arrangement near the crystal edge. Additional data may be obtained when SiPM detectors are connected to the sides 47 and/or back plane 48 of the crystal 30, since Anger logic can then be applied to even the outer most SiPMs 20. Since the SiPMs are only approximately .5 mm thick, the addition of SiPMs on the sides 47 and/or back plane 48 do not substantially alter the overall size of the detector arrangement, which is substantially smaller than the conventional PMT arrangement shown in Figure 1 .
- a conventional detector arrangement is approximately 275 mm thick: approximately 250 mm for the photo multiplier tube PMT array, approximately 16 mm for the light guide LC, and approximately 10 mm for the scintillator plate X.
- the embodiment shown in Figure 2 is approximately 16 mm thick: approximately 1 mm for two layers of SiPMs 20, approximately 5 mm for two layers of light guides 35, and approximately 10 mm for the scintillator plate 30.
- the overall thickness of the conventional detector arrangement and the detector arrangement shown in Figure 2 can be modified depending on the type of scintillator used, the desired thickness of the light guides, the number of layers used for the light guides and SiPMs, and the type of SiPM, or other avalanche diode, used.
- the relative size reduction achieved by the present invention remains substantial. By reducing the overall thickness of the detector arrangement by a factor of 20 to 30, the overall size and weight that an imaging system gantry must support is substantially reduced.
- the size and the detection efficiency of the SiPMs 20 determine the spatial resolution of the detector.
- the spatial resolution of a detector with 25 percent detector efficiency and SiPMs 10 mm or under is approximately 1 .5 mm.
- the simulation shows that SiPMs of 10 mm or under generally have the same spatial resolution, while SiPMs larger than 10 mm have lower resolution (higher numerical values).
- the spatial resolution of the detector may also increase to 1 mm or less if a higher detection efficiency is achieved.
- FIG 4 illustrates another embodiment 10' of the present invention.
- the components of the detector arrangement are the same as shown in Figure 2, however, the scintillator plate 30 is replaced with scintillator pixels 50.
- the scintillator pixels 50 are directly coupled to the SiPMs 20 in a one-to-one correspondence. In such embodiments, light guides are not generally required.
- the detector arrangement 10 described herein further provides a depth-of-interaction (DOI) measurement, which is enabled by the better sampling available due to the relatively small pixels.
- DOI depth-of-interaction
- the enhanced DOI measurement is especially valuable for providing better spatial resolution under oblique incidence.
- the use of a layer of SiPMs 20 on the back plane 48 of the crystal 30 provides even further DOI information.
- Figures 6A and 6B illustrate one implementation of the detector arrangement described herein.
- Figure 5 illustrates a prior art version of a detector 100.
- the arrangement of the prior art detector 100 creates a detector area 1 10 that is less than the camera area 1 1 5. Consequently, when two cameras are placed side-by-side, as shown in Figure 5, there are areas that are non-imaged 1 20, including a non-imaged area between the two detector areas 1 10.
- a complete image area is created. This is because the detector area 1 50 is equal to, or nearly equal to, the camera area 160. This leaves non-imaged areas 165 that are negligible.
- larger complete imaging areas can be important when viewing larger patients or larger portions of patients, such as full body scans.
- the more complete imaging area allows for simultaneous imaging of a larger imaged area, thereby increasing overall scan time.
- FIG. 8A in comparing the prior art detector arrangement 100 shown in Figure 8A with the embodiment of the new detector arrangement 140 shown in Figure 8B, it shown how placement of multiple cameras employing the new detector arrangement 140 provides more complete imaging of the imaged area.
- portions of the patient 1 80 lie outside the detector area 1 1 0, thus producing an incomplete image.
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- Physics & Mathematics (AREA)
- Health & Medical Sciences (AREA)
- Engineering & Computer Science (AREA)
- Biomedical Technology (AREA)
- General Health & Medical Sciences (AREA)
- Medical Informatics (AREA)
- Nuclear Medicine, Radiotherapy & Molecular Imaging (AREA)
- Optics & Photonics (AREA)
- Life Sciences & Earth Sciences (AREA)
- General Physics & Mathematics (AREA)
- High Energy & Nuclear Physics (AREA)
- Molecular Biology (AREA)
- Spectroscopy & Molecular Physics (AREA)
- Measurement Of Radiation (AREA)
- Nuclear Medicine (AREA)
- Luminescent Compositions (AREA)
- Apparatus For Radiation Diagnosis (AREA)
Abstract
Description
Claims
Priority Applications (5)
Application Number | Priority Date | Filing Date | Title |
---|---|---|---|
AT06780232T ATE537466T1 (en) | 2005-08-26 | 2006-07-27 | HIGH RESOLUTION MEDICAL IMAGING DETECTOR |
JP2008527540A JP2009506316A (en) | 2005-08-26 | 2006-07-27 | High resolution medical imaging detector |
US12/063,769 US8884239B2 (en) | 2005-08-26 | 2006-07-27 | High resolution medical imaging detector |
EP06780232A EP1922564B1 (en) | 2005-08-26 | 2006-07-27 | High resolution medical imaging detector |
CN200680031135.8A CN101248370B (en) | 2005-08-26 | 2006-07-27 | High resolution medical imaging detector |
Applications Claiming Priority (2)
Application Number | Priority Date | Filing Date | Title |
---|---|---|---|
US59604705P | 2005-08-26 | 2005-08-26 | |
US60/596,047 | 2005-08-26 |
Publications (1)
Publication Number | Publication Date |
---|---|
WO2007023401A1 true WO2007023401A1 (en) | 2007-03-01 |
Family
ID=37561256
Family Applications (1)
Application Number | Title | Priority Date | Filing Date |
---|---|---|---|
PCT/IB2006/052584 WO2007023401A1 (en) | 2005-08-26 | 2006-07-27 | High resolution medical imaging detector |
Country Status (7)
Country | Link |
---|---|
US (1) | US8884239B2 (en) |
EP (1) | EP1922564B1 (en) |
JP (1) | JP2009506316A (en) |
CN (1) | CN101248370B (en) |
AT (1) | ATE537466T1 (en) |
RU (1) | RU2401440C2 (en) |
WO (1) | WO2007023401A1 (en) |
Cited By (4)
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FR2925698A1 (en) * | 2007-12-19 | 2009-06-26 | Chu | Positron emission tomography device i.e. tomograph for radioisotopic functional imaging utilized in nuclear medicine, has annular chamber that is constituted by crystals for converting light energy of gamma photons positron annihilation |
US7652257B2 (en) * | 2007-06-15 | 2010-01-26 | General Electric Company | Structure of a solid state photomultiplier |
US8063377B2 (en) | 2008-08-15 | 2011-11-22 | Koninklijke Philips Electronics N.V. | Crystal identification for high resolution nuclear imaging |
US10473797B2 (en) | 2013-12-23 | 2019-11-12 | Johnson Matthey Public Limited Company | Radiation detection apparatus and method |
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US9588230B2 (en) * | 2008-09-15 | 2017-03-07 | Siemens Medical Solutions Usa, Inc. | Systems and methods for calibrating a silicon photomultiplier-based positron emission tomography system |
US9040924B2 (en) | 2009-10-27 | 2015-05-26 | University Of Washington Through Its Center For Commercialization | Optical-interface patterning for radiation detector crystals |
WO2011121707A1 (en) * | 2010-03-29 | 2011-10-06 | 独立行政法人放射線医学総合研究所 | Three dimensional radiation position detector and method of identifying the detection position |
KR101175697B1 (en) | 2010-07-02 | 2012-08-21 | 서강대학교산학협력단 | Method of improving LCE and linearity of relationship between gamma-ray's energy and the number of photons impinged on photosensor array in PET module |
CN102735350A (en) * | 2011-04-08 | 2012-10-17 | 北京师范大学 | Silicon photo-multiplier structure, production and usage |
US9176241B2 (en) | 2011-08-03 | 2015-11-03 | Koninklijke Philips N.V. | Position-sensitive readout modes for digital silicon photomultiplier arrays |
JP2014210047A (en) * | 2013-04-18 | 2014-11-13 | 株式会社東芝 | X-ray ct apparatus |
CN105190358B (en) * | 2013-05-10 | 2019-08-30 | 皇家飞利浦有限公司 | Large-area flicker volume elements part and radiation detector and the radiation absorption event positioning system for using it |
CN105425270B (en) * | 2014-05-28 | 2020-06-12 | 上海联影医疗科技有限公司 | PET detector, and PET detector setting method and PET detector detection method |
US20160231439A1 (en) * | 2015-02-06 | 2016-08-11 | Thermo Fisher Scientific Messtechnik Gmbh | Device and method for detection of radioactive radiation |
US9696439B2 (en) | 2015-08-10 | 2017-07-04 | Shanghai United Imaging Healthcare Co., Ltd. | Apparatus and method for PET detector |
WO2019000389A1 (en) | 2017-06-30 | 2019-01-03 | Shanghai United Imaging Healthcare Co., Ltd. | System and method for positron emission tomography |
CN107942367A (en) * | 2017-11-24 | 2018-04-20 | 合肥吾法自然智能科技有限公司 | A kind of new γ photon high spatial resolution detection devices |
CN109459783B (en) * | 2018-09-30 | 2023-04-11 | 中派科技(深圳)有限责任公司 | PET device, multilayer crystal PET detector, electronic readout module and method thereof |
KR102103577B1 (en) * | 2019-08-09 | 2020-04-22 | 경북대학교 산학협력단 | Photo sensor |
US12013503B2 (en) * | 2022-10-07 | 2024-06-18 | Cintilight, Llc | Lateral crystal photodiode readouts and switched diode networks for processing nuclear events |
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- 2006-07-27 WO PCT/IB2006/052584 patent/WO2007023401A1/en active Application Filing
- 2006-07-27 EP EP06780232A patent/EP1922564B1/en active Active
- 2006-07-27 AT AT06780232T patent/ATE537466T1/en active
- 2006-07-27 JP JP2008527540A patent/JP2009506316A/en active Pending
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Cited By (4)
Publication number | Priority date | Publication date | Assignee | Title |
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US7652257B2 (en) * | 2007-06-15 | 2010-01-26 | General Electric Company | Structure of a solid state photomultiplier |
FR2925698A1 (en) * | 2007-12-19 | 2009-06-26 | Chu | Positron emission tomography device i.e. tomograph for radioisotopic functional imaging utilized in nuclear medicine, has annular chamber that is constituted by crystals for converting light energy of gamma photons positron annihilation |
US8063377B2 (en) | 2008-08-15 | 2011-11-22 | Koninklijke Philips Electronics N.V. | Crystal identification for high resolution nuclear imaging |
US10473797B2 (en) | 2013-12-23 | 2019-11-12 | Johnson Matthey Public Limited Company | Radiation detection apparatus and method |
Also Published As
Publication number | Publication date |
---|---|
US8884239B2 (en) | 2014-11-11 |
EP1922564A1 (en) | 2008-05-21 |
RU2401440C2 (en) | 2010-10-10 |
US20100176301A1 (en) | 2010-07-15 |
CN101248370B (en) | 2012-07-11 |
EP1922564B1 (en) | 2011-12-14 |
ATE537466T1 (en) | 2011-12-15 |
JP2009506316A (en) | 2009-02-12 |
RU2008111491A (en) | 2009-10-10 |
CN101248370A (en) | 2008-08-20 |
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